The present invention relates generally to nuclear magnetic resonance (NMR) imaging apparatus and, more specifically, to high performance receiver coils for NMR imaging of cardiac and thoraco-abdominal regions of a human body. The invention is particular relates to RF receiver coils which permit high quality imaging of the human heart.
Nuclear magnetic resonance (NMR) imaging is a known technique for obtaining cross sectional images through desired portions of a human body without exposing a patient to ionizing radiation. Briefly, the patient is placed in a static magnetic field which causes magnetic dipoles of atomic nuclei with spin (for example, hydrogen nuclei) to orient themselves with the magnetic lines of force. Like a spinning gyroscope, the spinning nuclei tend to precess with a certain angular frequency, known as the Larmor frequency. By means of a transmitter coil a radio frequency pulse is applied at a frequency which matches the natural precessional frequency. This causes the magnetic dipoles to quickly precess, while absorbing energy. When the excitation pulse ends, the nuclei briefly emit an RF signal known as the free induction signal or free induction decay. The magnetization vector of the nuclei eventually returns to its original position. This emitted RF energy can be detected by inductive coupling to a receiver coil and analyzed in view of the nature of the excitation pulse to build a set of data from which images can be constructed.
It will be understood and appreciated that the present invention is not directed to any of the various techniques for defining the cross sections for which images are constructed (e.g. phase encoding) and the techniques for actually constructing the images. Rather, the present invention assumes the existence of these known techniques. In general, the present invention is concerned with efficiently receiving the emitted RF energy from a region of the body in bulk to provide a signal suitable for analysis.
It is possible for a single coil to serve the transmitter and receiver functions. However, for optimum performance it is desirable to provide separate transmitter and receiver coils due to different design considerations.
For example, the transmitter coil should be large, have good RF homogeneity, and relatively low Q to provide broadband excitation.
The receiver coil, however, should be as small as possible consistent with the region of interest, have a reasonable amount of RF homogeneity, have a high Q (narrow bandwidth), be minimally sensitive to dielectric loading, be comfortable for the patient, and be easy to use so as to provide the minimum connection and set-up time. An RF receiver coil should maximize flux coupling from the patient to the coil surface, and at the same time, in order to prevent damage to the associated preamplifier, be reasonably orthogonal to the transmitter coil so as to minimize coupling between the two coils, and should also function in a plane perpendicular to the main (static) magnetic field.
One important performance criterion of a tuned receiver coil is the signal-to-noise ratio (SNR) which can be obtained. In theory, SNR can be calculated by formula based on the frequency, the magnetic field due to a unit current in the coil, sample volume, loaded circuit Q, filling factor, and resistive loss due to coil impedance. The first three of these variables are easily determined, but the others are not.
Theoretical analysis is limited by several factors. First, it is difficult to mathematically determine the ideal coil geometry for each organ of interest, giving rise to a large number of possible shapes. Even simple application of Faraday's law of induction may be inappropriate. For example, a multiple turn solenoid may be no better than a single turn solenoid of the same diameter when loaded with a biological sample. Electrical performance criteria such as coil Q may become irrelevant when comparing different coil geometries since Q is simply the ratio of inductive reactance to coil resistance, and in appropriate geometries may have good Qs.
The magnetic field strength of an NMR system is a fundamental determinant of magnetic resonance signal intensity. Although capable of providing diagnostic data for a broad spectrum of disease states, resistive magnet systems have suffered reduced signal-to-noise ratio (SNR) due to magnetic field strength limitations. Improved SNR can be obtained in these systems by reducing machine noise and optimizing the receiver portion of the system. Unique coil designs have facilitated high resolution studies of the central nervous system, extremities, and head and neck regions. The use of various unique coils has allowed narrowing of slices and decreased pixel size using higher field gradients. The size and location of such regions facilitates imaging by conventional surface coils, as well as full and half-saddle coils.
Imaging of cardiac and thoraco-abdominal areas has however usually depended on the use of volume coils (filling factor &gt;50%), relatively large in size. Previous volume coils could not equal the spatial resolution possible using the surface coils (filling factor &lt;50%) because of inherently greater signal loss from large coils. Large coils suffer loss of signal because they require greater lengths of conductor, and tend to couple to their environment (gradient coils, the magnet assembly, and the like). The most severe loss is incurred by the presence of the whole human torso, which may be approximated by a large conductive cylinder. Eddy currents in the body absorb energy from the resonant circuit, resulting in a less efficient receiver coil (lower Q). This leads to reduced currents in the coil and a lower SNR. Further, with large coils there is a phenomenon of spatial aliasing seen along the phase encoding direction when image field of view is reduced.
Flat surface coils have been used to improve resolution of abdominal structures; however, some of the images produced have been limited due to poor deep penetration which precluded optimal evaluation of all tissue within the slice. These flat surface coils were often used after volume coils had screened the area at lower levels of resolution. Thus, imaging time was increased due to repeat acquisition and coil changing time.
In summary, during the course of NMR imaging, energy is inductively coupled from the patient to an RF receiver coil, where the emf induced in the coil and therefore the SNR available for image construction depends in part on flux coupling between the patient and the coil. Current NMR imaging systems are limited to thick slice (1 cm) medium resolution images of the body due in part to rf coil inefficiency. The SNR requirements for thin section (0.6 cm) images with increased in-plane resolution may require more than 100% greater SNR to obtain images with useful diagnostic content. Numerous RF coil designs have been suggested in the literature including saddle coils, half saddles, birdcage resonators and various distributed phase designs. Due to the size of the human body these coils have tended to be very large. Unfortunately, large coils do not have high flux coupling to any particular area of interest. Instead they have a moderate coupling to the entire body. While this may be acceptable for randomly imaging the body, it is a very inefficient means for imaging a particular organ of interest. Thus large body rf coils are available for NMR imaging of the entire torso which lack thin slice, high resolution capability in general.